Gamma ray detectors are used in several different applications, including in a positron emission tomography (PET) apparatus. PET is a nuclear medicine imaging technique that produces a three-dimensional image or picture of functional processes within a body. The system detects pairs of gamma rays that are emitted indirectly by a positron-emitting radionuclide (tracer) that is introduced into the body on a biologically active molecule. Images of the tracer concentration in 3-dimensional space within the body are reconstructed by computer analysis. In modern scanners, this reconstruction is often accomplished with the aid of a CT x-ray scan performed on the patient during the same session and using the same apparatus.
To begin the PET imaging process, a short-lived radioactive tracer isotope is incorporated into a biologicaly active molecule and injected into a patient (usually into the blood stream). After a short waiting period, the active molecule concentrates in the tissues of interest and the patient is placed in an imaging scanner. The molecule most commonly used for this purpose is fluorodeoxyglucose (FDG), a sugar labeled with an F-18 isotope with a half life of about 110 minutes.
As the radioisotope undergoes positron emission decay it emits a positron, which is the antiparticle of the electron. After traveling up to a few millimeters within the patient's body, the positron encounters an electron. The encounter annihilates both the positron and the electron, producing a pair of annihilation (gamma) photons that move in opposite directions, i.e., away from each other. The gamma rays are detected when they reach a scintillator in the scanning device, creating a burst of light that is detected by a photo-sensor (e.g., a photomultiplier tube (PMT), a silicon avalanche photodiode or a solid state photomultiplier (SSPM)). Detection depends on simultaneous or coincident detection of the two back-to-back photons, each photon detected by one of two detectors, with the two detectors placed in opposite directions from the annihilation location. Photons that do not arrive in temporal “pairs” (i.e. within a timing-window of few nanoseconds (ns), e.g., less than about 7 ns) are rejected/discarded by the scanner. State-of-the-art scanners are capable of determining the difference in arrival time of the annihilation photons to within about 0.5 ns at full width at half maximum (FWHM).
The radionuclides used in PET scanning are typically isotopes with short half-lives of less than about two hours. For example, an O-15 isotope has a half-life of about 123 seconds and the F-18 isotope referenced above has a half-life of about 110 minutes. These radionuclides are incorporated either into compounds normally used by the body such as glucose (or glucose analogues), water or ammonia, or into molecules that bind to receptors or other sites of physiological significance. Such labeled compounds are known as radiotracers. PET technology can be used to trace the biologic pathway of any compound in living humans provided such compound can be radio-labeled with a PET isotope. When the FDG molecule, an analogue of glucose, is used as the carrier, the imaged tracer concentrations provide information about tissue metabolic activity related to regional glucose uptake. Although FDG is the tracer most commonly used for clinical PET scans, other tracer molecules are used in PET devices to image the tissue concentration of many other molecules of interest.
The PET imaging system comprises, in one embodiment, a plurality of detector rings arranged coaxially to form a cylinder. After receiving the tracer isotope, the patient placed in the cylinder to detect the gamma rays emitted during the annihilation events. FIG. 1 illustrates one such detector ring 10 and a mass or body 12 placed within the ring 10. The ring 10 comprises a plurality of detector blocks 16 that detect gamma rays 20 emitted during annihilation events. The gamma rays travel in opposite directions and strike two oppositely-disposed detectors 20.
FIG. 2 illustrates a perspective view of a plurality of detector rings 10 and detector blocks 16.
FIG. 3 illustrates details of a single block detector 16 comprising a plurality of scintillation crystals 30 that are struck by the gamma rays. A plurality of photomultiplier tubes 34 (vacuum photomultiplier tubes or PMTs) are coupled to the scintillation crystals 30 to detect the light emitted during the scintillation process. FIG. 3 illustrates a 2×2 array of PMTs. A typical block detector comprises an M×N crystal array (with M and N greater than about 6) and a 2×2 array of PMTs (that is, four PMTs). Typically, 40-64 crystals are coupled to the four PMTs.
A one-to-one coupling of the crystals and the PMTs is not possible due to a thickness of the PMT glass. Further, manufacture of a one-to-one coupling block detector is expensive. Thus to reduce the cost and complexity, the 2×2 PMT array determines the incident gamma ray energy and also identifies the crystal in the array that received the gamma ray energy.
Energy deposited in the block detector of FIG. 3 (i.e., energy incident on any of the M×N crystals) is determined and readout by each of the four PMTs in the 2×2 PMT array. The energy in the four PMTs is combined to determine the total incident energy. As is known by those skilled in the art, Anger logic is used to determine the specific crystal that was struck by the incident gamma ray.
Detection by the ring of detectors (as in FIGS. 2 and 3) is based on the concept that two photons detected in close temporal proximity (e.g., within less than about 7 nanoseconds by the two oppositely disposed detectors) are likely to have originated from a single annihilation event in the patient's body somewhere along a line that connects the two detectors. Such a simultaneous detection is termed a “coincidence.” All of the coincidence events detected during an imaging session are recorded by the PET scanner as raw data. As in a single photon emission computed tomography (SPECT), the coincidence data in PET imaging is reconstructed by a computer to produce a three-dimensional image volume.
The electron-positron decays cause the emission of two 511 keV gamma photons at almost 180 degrees apart; hence it is possible to localize their source along a straight line of coincidence (also referred to as a line of response or LOR) connecting the two detected gamma photons. In practice, the LOR has a finite width as the emitted photons are not exactly 180 degrees apart.
If the resolving time of the detectors is greater than about 1 ns, it is difficult to localize the origin of the gamma rays to a segment of the LOR. If the timing resolution is better than about 1 ns, the event can be localized to a segment of the LOR. This localization process is referred to as time-of-flight detection and is used by modern systems with a high timing resolution that can precisely determine the time difference between detection of the photons. These systems thus reduce the length of the LOR segment of interest and more precisely determine the location of the origin of the gamma ray. As the timing resolution improves, the signal-to-noise ratio (SNR) of the reconstructed image also improves, requiring fewer events to achieve the same image quality.
The raw data collected by a PET scanner comprise a list of coincidence events representing near-simultaneous detection of annihilation photons by the pair of oppositely disposed detectors. Using statistics collected from hundred thousands of coincidence events, the most likely activity distribution can be computed using iterative reconstruction techniques known in the art, and thus a map of radioactivities, as a function of voxel (volume element) location parcels is constructed and displayed. The resulting map shows the tissues in which the molecular probe has become concentrated and this map can be interpreted by a nuclear medicine physician or radiologist.
Since the gamma rays are emitted from within the tissue, photon attenuation and absorption in bodily tissue between the annihilation site and the detectors may result in only one of the two photons reaching a detector. These are referred to as “single events” and the data associated with any such single events are discarded. The detection of more coincident events leads to improved sensitivity and resolution of the final image.
A typical PET detector employs a scintillation crystal area of about 4×4 cm2. The crystal area comprises a plurality of crystals, and thus is also referred to as a crystal array. The PET detector further comprises four PMTs (arranged in a 2×2 array), each PMT generating one detector signal. Each PMT is one element of a readout channel. Thus four detector signals (and thus four readout channels) cover the 4×4 cm2 crystal array. A high-gain high-bandwidth amplifier following each PMT amplifies the PMT output signal for input to additional readout/display components.
Photodiodes and solid state photomultipliers (SSPMs) can be used in lieu of the PMTs to detect the light emitted by the scintillation crystals. Since the photodiodes and SSPMs are smaller than the PMTs, they can accommodate a one-to-one coupling with the scintillation crystals. Also, the crystals utilized in an SSPM detector can be smaller than the crystals used with large PMTs, since with direct coupling there is no need to decode the signals from a few PMTs into many crystals; instead there is direct correspondence between the SSPM detector that detects a scintillation, and the crystal in which this scintillation occurred. On the other hand, because SSPMs are smaller than PMTs, many more SSPMs are needed to cover the same detector area. For example, covering a scintillation crystal area of about 4×4 cm2 may require as many as 100 SSPMs, compared with four PMTs.
The additional SSPMs and the one-to-one coupling with the crystals provide considerably improved timing and spatial resolution, but also create problems due to the requirement for higher density in the processing electronics components, resulting in the dissipation of additional power within a smaller space. The smaller area occupied by each crystal may also lead to increased spreading of the incident energy due to scattering (referred to as Compton scattering) from the crystal struck by the incident beam to proximate crystals. Thus although the use of smaller crystals provides better spatial resolution, it also increases the probability of Compton scattering, which leads to a decrease in timing resolution for the crystal signals.
FIG. 4 illustrates a plurality of crystals 48 represented as forming a checkerboard pattern of crystal elements to aid with the description of the invention. As is known by those skilled in the art, each crystal in the array is usually rectangular or square in shape such that a compact array can be formed, although other shapes (for example, hexagonal or triangular) can also be employed to form a compact array.
A gamma ray 50 (an annihilation photon) strikes a crystal 52 with no scattering. i.e., the ray deposits all its energy in the crystal 52. A gamma ray 56 strikes an electron within the crystal 58 and deposits a fraction of its energy there, then impinges a proximate crystal 62 where it deposits the remainder of its energy. This process is referred to as a Compton scattering event. In the latter case the gamma ray energy is absorbed in both the crystal 58 and the crystal 62. Given the conservation of energy principle, the sum of these energies equals the energy in the incident gamma ray. The relative values of these two energies depends on the nature of the collision, the energy of the incident gamma ray, and the mass of the particle struck (an electron) by the incident gamma ray in the crystal 58. Crystal strikes occurring within a predetermined time interval in proximate crystals are assumed to be from the same initial gamma ray. The two strikes will be separated by the time required for the scattered gamma ray to travel from the first crystal to the second crystal.
Two common circuits used to determine the timing of crystal strikes are the leading edge discriminator and the constant fraction discriminator. The leading edge discriminator is a much simpler circuit than the constant fraction discriminator, but the time measured by the leading edge discriminator is dependent on the amplitude of the signal resulting from a strike. The constant fraction discriminator measurement is independent of the signal amplitude. This amplitude dependence is referred to as walk. The energy deposited in multiple crystals from a gamma ray that scattered within the crystal array is less than the initial energy of the gamma ray.
In a one-to-one coupled detector, the walk causes the timing measurements of the strikes in the crystals for a Compton scattering event to be different than the timing that would have been measured if all the gamma ray energy had been deposited in a single crystal. If the signal amplitudes for all the strikes are measured, a correction can be made to eliminate any error in the timing of the strikes due to walk. It should be noted that in a PET scanner using PMT detectors, the timing is measured from the sum of the signals from all crystals and therefore does not contain any errors from Compton scattering walk.
To perform Compton scattering walk correction for a 10×10 crystal-SSPM combination, one-hundred time-to-digital converters (TDCs) are required. The prior art crystal-SSPM combination also requires one-hundred analog-to-digital converters (ADCs) for converting the energy signal to digital values to perform walk correction and obtain accurate timing information from each event. Thus the prior art requires one hundred TDCs and one hundred ADCs to reduce the Compton scattering effects. Each readout channel therefore comprises one crystal-SSPM combination, one TDC and one ADC.
Disadvantageously, using a TDC and an ADC in each readout channel requires an excessive amount of power and for this and other reasons, is not a practical solution. This prior art implementation also requires a substantial space for one hundred TDCs and one hundred ADCs.
To perform the Compton scattering correction, it is necessary to determine for each crystal in which the gamma ray interacted, how much energy was deposited in that crystal and the time that the energy was deposited in the crystal. However, this determination must be made with a limited number of components due to physical space limitations. It is also desired to limit the power consumption/dissipation of these components. The present invention discloses a scheme for satisfying these constraints.